The invention relates to a multi-channel switching system for a multi-channel gradient coil system for MRI (=Magnetic Resonance Imaging), comprising: a plurality of Nswitch analog switches to connect a plurality of Nelement coil elements, whereby said plurality of analog switches and coil elements forms a plurality of Nchannel electrical channels each driven by a gradient power amplifier; a distribution board to generate control signals for each of the plurality of analog switches; a digital controller providing the command code to the distribution board through a communication bus; and a power delivery system to power each of Nswitch analog switches.
A multi-channel switching system of this type for shimming is known e.g. from Harris et al., “A new approach to shimming: the dynamically controlled adaptive current network”, MRM 71: 859-869, 2014. Such systems, however, require individual power supplies for each switch, which limits the scalability.
The present invention relates generally to magnetic resonance imaging (MRI). It specifically relates to shimming and spatial encoding hardware for MRI.
Magnetic resonance imaging (MRI) is a relative new technology compared with computed tomography (CT) and the first MR Image was published in 1973 by P. C. Lauterbur in “Image Formation by Induced Local Interactions: Examples of Employing Nuclear Magnetic Resonance”, Nature 242, 190491. It is primarily a medical imaging technique which most commonly used in radiology to visualize the structure and function of the body. It could provide detailed Images of the body in any plane. MRI provides much greater contrast between the different soft tissues of the body than CT does, making it especially useful in neurological, cardiovascular, and oncological imaging. It uses a powerful magnetic field to align the nuclear magnetization of hydrogen atoms in water in the body. Radio frequency fields are used to systematically alter the alignment of this magnetization, causing the hydrogen nuclei to produce a rotating magnetic field detectable by the scanner. This signal can be manipulated by additional magnetic fields to build up enough information to reconstruct an image of the body.
An MRI system typically establishes a homogenous magnetic field, generally along a central axis of a subject undergoing an MRI procedure. This homogenous main magnetic field affects the magnetic properties of the subject to be imaged by aligning the nuclear spins, in atoms and molecules forming the body tissue. If the orientation of the nuclear spins is perturbed out of alignment, the nuclei attempt to realign their spins with the field. Perturbation of the orientation of the nuclear spins is typically caused by application of radio frequency (RF) pulses tuned to the Larmor frequency of the material of interest. During the realignment process, the nuclei precess about the direction of the main magnetic field and emit electromagnetic signals that may be detected by one or more RF detector coils placed on or around the subject.
Magnetic resonance imaging employs temporally and spatially variable magnetic fields to encode position by affecting the local Larmor frequency of spins. Gradient coils typically used for that purpose generate spatial encoding magnetic fields (=SEMs) which are superimposed on the main magnetic field. This allows to choose the localization of the image slices and also to provide phase encoding and frequency encoding. This encoding permits identification of the origin of resonance signals during image reconstruction. The image quality and resolution depends significantly on the strength and how the applied encoding fields can be controlled. Control of the gradient coils is generally performed in accordance with pre-established protocols or sequences at events, called pulse sequences, permitting different types of contrast mechanisms to be imaged.
Gradient coils are typically designed to generate spatial encoding magnetic fields in a linear fashion, i.e. constant linear gradient fields, along the three orthogonal directions X, Y and Z. Typically, a gradient coil operates with maximum currents of about few hundreds amperes and at maximum voltages in a range from about few hundreds volts to about few thousands volts. To achieve higher resolution of an image, stronger and faster gradient fields are needed. Therefore gradient coils need higher currents and voltages. However, this induces safety concerns due to peripheral nerve stimulation (=PNS) and increases the complexity and cost of the current sources which are referred to as gradient power amplifiers. Another drawback of linear gradients is the missing flexibility in terms of realizable field shapes.
To overcome these limitations non-linearly varying SEMs have been introduced by EP 1 780 556 B1. Such gradient fields may overcome PNS limits and allow for parallel non-bijective image encoding or curved slice imaging (see e.g. EP 2 511 725 A1). For faster image acquisition more encoding fields than spatial dimensions (X, Y and Z) have been used (for example 4D-Rio, O-Space, see e.g. US 2012/0 286 783 A1). Multi element coil systems have been used to generate the SEMs needed for those applications. However, such systems still have a limited set of field shapes and need a dedicated gradient power amplifier per coil element.
Susceptibility differences in the object to be imaged introduce field perturbations of the homogeneous main magnetic. Compensation of these perturbations is commonly referred to as shimming. Shim coils used for these corrections usually aim to generate shim fields that are based on spherical volume harmonics. However, complex susceptibility differences present in the human body introduce perturbations that cannot be fully corrected for by spherical volume harmonics of lower orders. To overcome these limitations a set of multiple self-similar individual coil elements, each supplied by a separate current source has been introduced (by Christoph Juchem et al in “Magnetic field homogenization of the human prefrontal cortex with a set of localized electrical coils”, MRM 63:171-180, 2010). Although this approach gains flexibility in terms of realizable field shapes used for shimming, the number of individual current sources scales linearly with the number of coils. This increases cost and space needed. To significantly reduce the number of amplifiers while allowing to route currents along pre calculated paths a Dynamically Controlled Adaptive Current Network has been introduced (see Chad. T. Harris et al. in “A new approach to shimming: the dynamically controlled adaptive current network”, MRM 71: 859-869, 2014).
The use of a multiple of individual coil elements to generate SEMs has been introduced for example in WO 2009/124873 A1. This allows for enhanced flexibility in terms of realizable field profiles of the SEMs. However, as with the set of localized electrical coils used for shimming the number of GPAs and the corresponding connection cables equals the number of individual coil elements.
Switches to route currents in gradient setups have been presented in U.S. Pat. No. 6,157,280. However, the proposed switching method is either too slow to be changed within a pulse sequence (fluid actuated switch) or does not teach us on how the scaling constraint of requiring a separate power supply for each switch can be overcome. Furthermore, this patent did not describe a way of generating different field shapes with the same coil set up.
The present invention presents a way to substantially overcome one or more disadvantages of the existing arrangements.
An object of the present invention therefore is to provide a multi-channel switching system with the features defined initially, which allows to electrically connect matrix coil elements dynamically within a pulse sequence to generate dynamically switched magnetic field profiles, and therefore reduce the number of gradient power amplifiers, gradient cables and power supplies needed.